Methods for pet detector afterglow management

ABSTRACT

Disclosed herein are methods and devices for the acquisition of positron emission (or PET) data in the presence of ionizing radiation that causes afterglow of PET detectors. In one variation, the method comprises adjusting a coincidence trigger threshold of the PET detectors during a therapy session. In one variation, the method comprises adjusting a gain factor used in positron emission data acquisition (e.g., a gain factor used to multiply and/or shift the output(s) of a PET detector(s)) during a therapy session. In some variations, a method for acquiring positron emission data during a radiation therapy session comprises suspending communication between the PET detectors and a signal processor of a controller for a predetermined period of time after a radiation pulse has been emitted by the linac.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims priority to U.S. Provisional Patent ApplicationNo. 62/531,260, filed Jul. 11, 2017, which is hereby incorporated byreference in its entirety.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

This invention was made in part during work supported by grant number2R44CA153466-02A1 from the National Cancer Institute. The government mayhave certain rights in the invention.

TECHNICAL FIELD

This disclosure relates to methods for use in a radiation therapy systemcomprising a linear accelerator (or other ionizing radiation source) andone or more positron emission (or PET) detectors.

BACKGROUND

Radiation therapy systems typically have a radiation source (e.g., alinear accelerator or linac) that generates therapeutic radiation beamsfor the irradiation of targeted tissue regions, such as patient tumorregions. Although the generated radiation beams may be directed towardtargeted regions and may be beam-limited by one or more jaws and/orcollimators, a portion of the radiation beams may deviate and/or scatterfrom the targeted regions. This scattered radiation may interfere withthe function of other components of the radiation therapy system.

For example, scattered or stray radiation may affect the ability ofvarious detectors in a radiation therapy system, such as X-ray and/orPET detectors, to precisely acquire data. PET detectors in a radiationtherapy system may be affected such that the PET detector response toscattered or stray radiation may be indistinguishable from true positronemission events. In situations with high levels of radiation (e.g.,during a radiation pulse from a linac, for example), the PET detectorsmay “blank” and/or saturate. This may render them incapable ofmeaningfully detecting positron emission data.

Accordingly, it may be desirable to develop methods and devices tomanage the risk of equipment damage and/or data corruption due toscattered radiation from the linac.

SUMMARY

Disclosed herein are methods and devices for the acquisition of positronemission (or PET) data in the presence of ionizing radiation that causesafterglow of PET detectors. In one variation, the method may compriseadjusting a coincidence trigger threshold of the PET detectors during atherapy session. The coincidence trigger threshold may be increased asthe degree of PET detector afterglow increases. For example, thecoincidence trigger threshold may be increased as a dark count rate ofone or more of the PET detectors increases and/or exceeds a thresholddark count rate. Alternatively or additionally, the coincidence triggerthreshold may be increased as a bias current of one or more of the PETdetectors increases and/or exceeds a threshold bias current level. Thecoincidence trigger threshold may also be adjusted based on a measuredtemperature of the system (e.g., at or around the PET detectors), wherethe coincidence trigger threshold may be increased as the temperature ofthe system increases. In some variations, the coincidence triggerthreshold may be adjusted based on the radiation output of the radiationsource or linac. For example, the coincidence trigger threshold may beadjusted when the number of emitted radiation pulses exceeds apredetermined threshold, and/or based on a pulse schedule, and/or basedon the cumulative amount of radiation emitted by the linac during atherapy session. In some variations, the coincidence trigger thresholdmay be adjusted if the synchronization between two system components(e.g., linac and collimator) shifts, and the timing shift exceeds apredetermined threshold.

In some variations, a method for acquiring positron emission data duringa radiation therapy session may comprise suspending communicationbetween the PET detectors and a signal processor of a controller for apredetermined period of time after a radiation pulse has been emitted bythe linac. For example, the predetermined period of time may be about100 μs or more, or about 200 μs or more. Alternatively or additionally,the predetermined period of time may be determined at least in part by awidth or duration of a linac radiation pulse. For example, thepredetermined period of time may be about 25 times or about 100 timeslonger than the duration of a linac pulse. After the predeterminedperiod of time has elapsed, communication between the PET detectors andthe signal processor may resume and positron emission data may betransmitted from the detectors to the signal processor and/or acquiredby the signal processor for analysis and/or storage by the controller.

In other variations, a radiation therapy system may comprise a radiationsource, a plurality of PET detectors (e.g., PET detector arrays), and aradiation-blocking shield movable over the plurality of PET detectors.The radiation-blocking shield may be positioned over the PET detectorsduring an irradiation interval when the radiation source is emittingradiation, and may be positioned away from the PET detectors during adetection interval when the radiation source is not emitting radiation.

One variation of a radiation therapy system may comprise a radiationsource configured to direct one or more radiation pulses toward aPET-avid region of interest, where each radiation pulse has apredetermined pulse duration, a plurality of PET detectors configured todetect a positron emission path by detecting a pair of positronannihilation photons incident upon a portion of the detectors within acoincidence time-window and that generate a detector signal that exceedsa coincidence trigger threshold, and a controller in communication withthe plurality of PET detectors, where the controller is configured toadjust the coincidence trigger threshold during a therapy session. Thecontroller may be configured to adjust the coincidence trigger thresholdafter a threshold number of radiation pulses have been directed towardthe region of interest. The threshold number of radiation pulses may beapproximately 1,000 radiation pulses. The coincidence trigger thresholdmay be from about two photon-triggers to about five photon-triggers. Thecoincidence trigger threshold may be a first coincidence triggerthreshold and the threshold number of radiation pulses may be a firstthreshold number of radiation pulses, and the controller may beconfigured to adjust the first coincidence trigger threshold to a secondcoincidence trigger threshold after a second threshold number ofradiation pulses have been directed toward the region of interest. Thesecond coincidence trigger threshold may be greater than the firstcoincidence trigger threshold and the second threshold number ofradiation pulses may be greater than the first threshold number ofradiation pulses. The second coincidence trigger threshold may be fromabout four photon-triggers to about six photon-triggers, and the secondthreshold number of radiation pulses may be about 2,000. The secondcoincidence trigger threshold may be less than the first coincidencetrigger threshold and the second threshold number of radiation pulsesmay be greater than the first threshold number of radiation pulses. Thecontroller may be configured to adjust the coincidence trigger thresholdbased on changes in timing greater than 10% from baseline and/or may beconfigured to adjust the coincidence trigger threshold when a dark countrate of one or more of the plurality of PET detectors exceeds athreshold dark count rate. The threshold dark count rate may be fromabout 3 Mcps to about 10 Mcps, for example. Optionally, the controllermay further comprise a current detector configured to measure a biascurrent of one or more of the plurality of PET detectors, and whereinthe controller is configured to adjust the coincidence trigger thresholdwhen the bias current exceeds a threshold bias current value. Thethreshold bias current value may be from about 0.1 mA to about 5 mA,e.g., about 1 mA, about 3 mA. Alternatively or additionally, thecontroller may be configured to adjust the coincidence trigger thresholdwhen the amount of radiation emitted from the radiation source exceeds athreshold radiation level. The threshold radiation level may be fromabout 0.1 cGy/min to about 1 cGy/min. The controller may furthercomprises a signal processor and a switch configured to selectivelycommunicate a PET detector output signal to the signal processor. Theswitch may be configured to suspend communication of the PET detectoroutput signal to the signal processor for a predetermined period of timefollowing each radiation pulse, where a ratio of the predeterminedperiod of time to the duration of each radiation pulse may be betweenabout 25:1 to about 100:1. The controller may be configured to suspendcommunication of the PET detector output signal to the signal processorfor the duration of each radiation pulse and the predetermined period oftime following each radiation pulse. The controller may be configured tosuspend communication of the PET detector output signal to the signalprocessor based on a gate signal. The gate signals may cause thecontroller to suspend communication of the PET detector output signal tothe signal processor for at least 100 μs following each radiation pulse.In some variations, the gate signal may cause the controller to suspendcommunication of the PET detector output signal to the signal processorfor at least 200 μs following each radiation pulse. Alternatively oradditionally, the controller may be configured to adjust the coincidencetrigger threshold at least partially based on a timing schedule of theradiation pulses.

Also disclosed herein is a method for automatically adjusting thecoincidence trigger threshold for PET detectors. The method may comprisemeasuring a characteristic of a radiation therapy system comprising twoor more PET detectors having a coincidence trigger threshold,determining whether the measured characteristic exceeds a pre-determinedthreshold for that characteristic, and adjusting the coincidence triggerthreshold based on the determination of whether the measuredcharacteristic exceeds the threshold for that characteristic. Adjustingthe coincidence trigger threshold may comprise increasing thecoincidence trigger threshold if the measured characteristic exceeds thepre-determined threshold for that characteristic or decreasing thecoincidence trigger threshold if the measured characteristic is at orbelow the pre-determined threshold for that characteristic. The measuredcharacteristic may be a dark count rate of the two or more PET detectorsand the pre-determined threshold may be a dark count rate threshold. Themeasured characteristic may be a bias current of the two or more PETdetectors and the pre-determined threshold may be a bias currentthreshold. The radiation therapy system may comprise a temperaturesensor, and the measured characteristic may be a temperature measurementand the pre-determined threshold may be a temperature threshold.Alternatively or additionally, the radiation therapy system may comprisea radiation source having a pulse counter, and the measuredcharacteristic may be a pulse count measured from the pulse counter andthe pre-determined threshold may be a pulse count threshold. Theradiation therapy system may comprises a radiation source and acollimator, where the radiation source and the collimator may beconfigured to operate together with a pre-determined timing tolerance,and where the measured characteristic may be the amount of deviationfrom the pre-determined timing tolerance and the pre-determinedthreshold may be a timing deviation threshold.

Also disclosed herein is a method for detecting positron annihilationemission paths. The method may comprise directing one or more radiationbeam pulses to a target region, where the target region is PET-avid,detecting a first positron emission path defined by a first pair ofpositron annihilation photons that are incident upon a portion of aplurality of PET detectors within a time-window and that generate adetector signal that exceeds a first coincidence trigger threshold,adjusting the first coincidence trigger threshold to a secondcoincidence trigger threshold, and detecting a second positron emissionpath defined by a second pair of positron annihilation photons that areincident upon a portion of the plurality of PET detectors within thetime-window and that generate a detector signal that exceeds the secondcoincidence trigger threshold. The first coincidence trigger thresholdmay be adjusted to a second coincidence trigger threshold after apredetermined number of radiation beam pulses have been directed to thetarget region. Adjusting the first coincidence trigger threshold may beat least partially based on a timing schedule of radiation pulses. Thesecond coincidence trigger threshold may have a greater value than thefirst coincidence trigger threshold, for example, the second coincidencetrigger threshold may be about four photon-triggers and the firstcoincidence trigger threshold may be about two photon-triggers. In somevariations, the predetermined number of radiation pulses may be about1,000. The predetermined number of radiation pulses may be a firstpredetermined number of radiation pulses and the method may furthercomprise adjusting the second coincidence trigger threshold to a thirdcoincidence trigger threshold after a second predetermined number ofradiation pulses have been directed to the target region and detecting athird positron emission path defined by a third pair of positronannihilation photons that are incident upon a portion of the pluralityof PET detectors within the time-window and that generate a detectorsignal that exceeds the third coincidence trigger threshold. The thirdcoincidence trigger threshold may be greater than the second coincidencetrigger threshold and the second predetermined number of radiationpulses may be greater than the first predetermined number of radiationpulses. The third coincidence trigger threshold may be from about fourphoton-triggers to about six photon-triggers, and the secondpredetermined number of radiation pulses may be about 2,000. Theradiation beam pulses may each have a pulse width, and the plurality ofPET detectors may be in communication with a controller comprising asignal processor, and the method may further comprise suspendingcommunication of data from the PET detectors to the signal processer isfor a predetermined period of time following each radiation pulse, wherea ratio of the predetermined period of time to the pulse width may bebetween about 25:1 and about 100:1. Optionally, suspending communicationof the data may be based on a gate signal. The gate signal may causesuspension of communication of data from the PET detectors to the signalprocessor for at least 100 μs following the radiation pulse, or the gatesignal may cause suspension of communication of data from the PETdetectors to the signal processor for at least 200 μs following eachradiation pulse. The first coincidence trigger threshold may be adjustedto a second coincidence trigger threshold when a dark count rate of oneor more of the plurality of PET detectors exceeds a threshold dark countrate. The threshold dark count rate may be from about 3 Mcps to about 10Mcps. The first coincidence trigger threshold may be adjusted to asecond coincidence trigger threshold when a bias current of one or moreof the plurality of PET detectors exceeds a threshold bias currentvalue. For example, the threshold bias current value may be from about0.1 mA to about 5 mA, e.g., about 1 mA, about 3 mA. The firstcoincidence trigger threshold may be adjusted to a second coincidencetrigger threshold when the amount of radiation emitted from theradiation source exceeds a threshold radiation level. For example, thethreshold radiation level may be from about 0.1 cGy/min to about 1cGy/min.

Also disclosed herein is a radiation therapy system comprising aradiation source configured to deliver one or more radiation pulsestoward a PET-avid region of interest during one or more irradiationintervals, a plurality of PET detectors configured to detect one or morepositron emission paths emitted by the PET-avid region of interestduring one or more detection intervals, and a radiation-blocking filtermovable over the plurality of PET detectors. The radiation-blockingfilter may be configured to be positioned over the plurality of PETdetectors during the one or more irradiation intervals and positionedaway from the PET detectors during the one or more detection intervals.

Disclosed herein is a radiation therapy system comprising a radiationsource configured to direct one or more radiation pulses toward aPET-avid region of interest a plurality of PET detectors that areconfigured to detect positron annihilation photons, a current detectorconfigured to measure a bias current of the plurality of PET detectors,and a controller configured to receive photon data output from theplurality of PET detectors, wherein the controller is configured todetect a pair of coincident positron annihilation photons by adjustingthe photon data output using a gain factor having a value that is basedon the measured bias current during a therapy session (e.g., calculatedbased on the measured bias current). The controller may be configured toadjust the gain factor when the bias current exceeds a threshold biascurrent value, e.g., the threshold bias current value may be from about0.1 mA to about 1 mA. In some variations, the gain factor may be a ratiobetween the measured bias current and a magnitude of a photopeak shiftof the detection of the positron annihilation photons in photon dataoutput. Adjusting the photon data output may comprise multiplying thephoton data output by the gain factor or linearly shifting the photondata output by the gain factor. Alternatively or additionally, thecontroller may be configured to adjust the gain factor after a thresholdnumber of radiation pulses have been directed toward the region ofinterest, e.g., the threshold number of radiation pulses may beapproximately 1,000 radiation pulses. In some variations, the gainfactor may be a first gain factor and the threshold number of radiationpulses may be a first threshold number of radiation pulses, and thecontroller may be configured to adjust the first gain factor to a secondgain factor after a second threshold number of radiation pulses havebeen directed toward the region of interest. The second gain factor maybe greater than the first gain factor and the second threshold number ofradiation pulses may be greater than the first threshold number ofradiation pulses. Alternatively or additionally, the controller may beconfigured to calculate a photopeak location of annihilation photonsbased on the photon data output from the plurality of PET detectors andto adjust the gain factor based on shifts of the photopeak location froma baseline level. Alternatively or additionally, the controller may beconfigured to adjust the gain factor when a dark count rate of one ormore of the plurality of PET detectors exceeds a threshold dark countrate, e.g., the threshold dark count rate is from about 3 Mcps to about10 Mcps. Alternatively or additionally, the controller may be configuredto adjust the gain factor when the amount of radiation emitted from theradiation source exceeds a threshold radiation level, e.g., thethreshold radiation level may be from about 0.1 cGy/min to about 1cGy/min.

In some variations, the controller may further comprise a signalprocessor and a switch configured to selectively communicate a PETdetector output signal to the signal processor. The switch may beconfigured to suspend communication of the PET detector output signal tothe signal processor for a predetermined period of time following eachradiation pulse, where a ratio of the predetermined period of time tothe duration of each radiation pulse may be between about 25:1 to about100:1. The controller may be configured to suspend communication of thePET detector output signal to the signal processor for the duration ofeach radiation pulse and the predetermined period of time following eachradiation pulse. For example, the controller may be configured tosuspend communication of the PET detector output signal to the signalprocessor based on a gate signal. In some variations, the gate signalmay cause the controller to suspend communication of the PET detectoroutput signal to the signal processor for 100 μs or more following eachradiation pulse, e.g., the gate signal may cause the controller tosuspend communication of the PET detector output signal to the signalprocessor for 200 μs or more following each radiation pulse.Alternatively or additionally, the controller may be configured toadjust the gain factor at least partially based on a timing schedule ofthe radiation pulses.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A is a schematic diagram of a front view of one variation of aradiation therapy system.

FIG. 1B schematically depicts scattered X-rays that may cause PETdetector afterglow and plots of PET detector outputs affected byafterglow.

FIG. 2A is a flowchart diagram of one variation of a method for dynamicgain adjustment.

FIG. 2B is a flowchart diagram of one variation of a method for dynamicPET detector threshold adjustment.

FIG. 3 is a flowchart diagram of one variation of a method for dynamicPET detector threshold adjustment based on PET detector noise levels.

FIG. 4 is a flowchart diagram of another variation of a method fordynamic PET detector threshold adjustment based on changes in componenttiming distributions.

FIG. 5 is a flowchart diagram of one variation of a method for dynamicPET detector threshold adjustment based on PET detector dark countrates.

FIG. 6A is a flowchart diagram of one variation of a method for dynamicPET detector threshold adjustment based on PET detector bias current.

FIG. 6B is a flowchart diagram of one variation of a method for dynamicgain adjustment based on PET detector bias current.

FIG. 7A is a flowchart diagram of one variation of a method for gatingthe communication of positron emission data from PET detectors to thecontroller.

FIG. 7B is a schematic diagram of one variation of a logic circuit forgating the communication of positron emission data from PET detectors tothe controller.

FIG. 7C is a timing diagram of one variation of a method for gating thecommunication of positron emission data from PET detectors to thecontroller.

FIG. 8A is a schematic depiction of one variation of a radiation filterring in a first configuration.

FIG. 8B is a schematic depiction of the radiation filter ring of FIG. 8Ain a second configuration.

FIG. 8C is a side-view of a schematic depiction of the radiation filterring of FIG. 8A in the first configuration.

FIG. 9A is a schematic depiction of another variation of a radiationfilter ring in a first configuration.

FIG. 9B is a schematic depiction of the radiation filter ring of FIG. 9Ain a second configuration.

FIG. 10A depicts the parameters and arrangement of an experimental setupfor measuring PET detector afterglow.

FIG. 10B is a schematic depiction of an arrangement of an experimentalsetup for measuring the dark count rate of single crystal PET detectors.

FIG. 10C provides plots of PET detector data and dark counts before,during, and after a linac pulse.

FIG. 10D provides plots of PET detector dark count rates after a linacpulse over time.

FIG. 10E provides plots of PET detector dark count rates after a linacpulse over time.

FIG. 11 depicts the parameters and arrangement of another experimentalsetup for measuring two coincidence multi-crystal PET detectorafterglow.

FIG. 12 is a plot of the time resolution of time-of-flight PET detectorsas a function of time after a linac pulse.

FIG. 13A depicts an experimental data plot of the changes in biascurrent and temperature as a function of linac beam on-time.

FIG. 13B depicts a calibration plot that has been generated by measuringthe bias current and photopeak location.

FIG. 13C depicts a plot of the energy resolution of a PET detector overtime (each data-series interval represents a 10 minute increment wherethe linac beam was turned on at data-series value 1 and turned off atdata-series value 7), with afterglow correction by gain adjustment.

FIG. 13D depicts the shift of a time resolution centroid over time (eachdata-series interval represents a 10 minute increment where the linacbeam was turned on at data-series value 1 and turned off at data-seriesvalue 7).

DETAILED DESCRIPTION

Some variations of radiation therapy systems may comprise a therapeuticradiation source (such as a linac) and one or more PET detectors (e.g.,one or more PET detector arrays) for detecting emissions from apositron-emitting (i.e., PET-avid) tissue region. A patient may beinjected with a molecule labeled with a radioactive atom, known as a PETradiotracer prior to a treatment session, and the tracer maypreferentially accumulate at one or more tumor regions. The radioactiveatoms inside the patient undergo radioactive decay and emit positrons.Once emitted from an atom, a positron will quickly collide with a nearbyelectron after which both will be annihilated. Two high energy photons(511 keV) are emitted from the point of annihilation and travel inopposite directions. When the two photons are simultaneously detected bytwo PET detectors, it is known that the annihilation occurred somewherealong the line joining the two PET detectors. Radiation therapy systemsmay acquire positron emission data before or during the treatmentsession, and this emission data may be used to guide irradiation ofthese tumor regions. For example, emission-guided radiation therapysystems may comprise a plurality of PET detectors and a linac that aremounted on a gantry that is rotatable about a patient. In somevariations, a plurality of PET detectors may comprise two PET detectorarrays that are disposed opposite each other on the gantry. Emissiondata acquired in real-time by the detectors may be analyzed by a systemcontroller to control the rotation of the gantry to direct radiationfrom the linac to the PET-avid tumor regions. In some variations,real-time positron emission data may also be used to update treatmentplans to account for any tumor movement that may have occurred betweenthe treatment planning session and the treatment session.

A PET detector comprises a scintillating material (e.g., a scintillatingcrystal such as bismuth germanium oxide, gadolinium oxyorthosilicate, orlutetium oxyorthosilicate), coupled to a sensor (e.g., anyphotodetector, a photomultiplier tube such as a siliconphotomultiplier). When a high-energy photon strikes a PET detector, theenergy from that photon causes a scintillation event in thescintillating material, which may generate one or more lower-energy(e.g., visible light) photons that are detected by the photodetectordevice. Photodetector devices may have a baseline dark count rate ordark current, where random fluctuations in the output may beindistinguishable from fluctuations that indicate the presence of aphoton. A dark count causes a pixel of a detector to fire bydischarging. When a pixel discharges, it draws current from the powersource, and the current drawn from the power source may be referred toas a bias current. The bias current may be proportional to the averagenumber of dark counts that have fired over a period time plus otherconstant or slowly varying terms; that is, the bias current may beproportional to the dark current. The dark current may be proportionalto the afterglow photocurrent plus the thermal noise current of the PETphotodetector. The bias current may be measured using acurrent-measurement device or module that may be included with a PETdetector array. Alternatively or additionally, the bias current may bemeasured using an ammeter disposed in series with the PET detectorphotodetector and the power source. Measuring the bias current and/orchanges to the bias current to the photodetector at a selected or setoperating range (e.g., gain and/or sensitivity) may provide anindication of the dark count rate and/or changes in the dark count rate(i.e., changes in the bias current may indicate shifts in the dark countrate). For example, as the dark count rate increases, the bias currentto the photodetector of the PET detector may also increase because morecurrent is drawn from the power source as a greater number of randomfluctuations causes a pixel of a detector to discharge more frequently.Under normal operating conditions, the dark count rate may be relativelylow, for example, approximately 2 million dark counts per second (cps).Increased ambient temperature and/or elevated levels of radiation maycause the dark count rate or dark current of a photodetector toincrease.

A radiation therapy system may comprise at least two arrays of PETdetectors located opposite each other on a gantry. For example, a PETdetector on a first array may have a corresponding PET detector on asecond array located on the opposite side so that the two high-energyphotons from a positron annihilation event may be detected. In onevariation, a radiation therapy system may comprise two PET detectorarrays, each comprising 32 PET detector modules (for a total of 64 PETdetector modules). Each PET detector module may comprise a 6×12 subarrayof PET detectors, where each PET detector has its own photodetector. Insome variations, each PET detector module may measure and output thebias current of all of the photodetectors in the 6×12 array of PETdetectors, and the gain of all of the photodetectors in the PET detectormodule may be set by a single gain input value. Since positron emissionand annihilation events are stochastic events, the PET detectors of asystem may detect a plurality of high-energy photons within a short timeinterval, and a controller uses the temporal information of eachdetected photon (e.g., time of detection), as well as the location ofthe PET detectors that detected these photons, to determine which twophotons are part of a positron annihilation photon pair. For example, iftwo high-energy photons are detected by two PET detectors that arelocated opposite to each other within a particular time interval (e.g.,a coincidence time-window), then the controller may pair these twophotons together as originating from the same positron annihilationevent, which occurred somewhere along the line joining these two PETdetectors. A coincidence time-window is the time interval within whichdetected photons may be considered coincident (and processed as if theyoriginate from the same positron annihilation event). The coincidencetrigger threshold may be a trigger threshold that discriminates betweensignals arising from the detection of an annihilation photon and signalsthat arise from scattered radiation and/or other noise sources (e.g.,random detector noise, afterglow, thermal noise, etc.). If the locationof the annihilation event is closer to one of the PET detectors than theother, one photon of the pair will have a shorter distance to travelthan the other (i.e., one photon will have a shorter time-of-flight thanthe other), and will therefore strike the first PET detector before thesecond photon strikes the second PET detector. The time differentialbetween the detection of the photons in a positron annihilation pair maybe used by the controller to determine where an annihilation eventoccurred on the line between the two PET detected events. PET detectorsthat have sufficient temporal precision to sense differences in thetime-of-flight (TOF) of positron annihilation photons may transmit TOFdata to a system controller for calculating the location of the positronannihilation event.

During a treatment session, the linac may generate pulses of high-fluxX-rays that are emitted toward the target regions. Beam-limitingdevices, such as one or more jaws and/or collimators (e.g., a multi-leafcollimator), may help to limit the spread of the X-rays and direct theX-rays to targeted tissue regions. These X-rays may interact with thepatient, where a portion of the X-rays irradiate the target regions inpatient (e.g., tumor regions), and a portion of the X-rays may bescattered by the patient. The scattered X-rays may interact withcomponents of the radiotherapy system, such as X-ray detectors (e.g., MVor kV detectors) and/or PET detectors. This effect is schematicallydepicted in FIG. 1A, where the body of a patient 120 may scatter X-raysfrom a linac 130 and target 132. The X-rays from the linac and targetmay be shaped by beam-limiting devices, such as a multi-leaf collimator134 to form a treatment beam 122. The scattered X-rays or radiation 124may be incident on the PET detectors 126, triggering scintillatingevents (e.g., lower-energy photons) that may be indistinguishable fromscintillating events caused by positron emissions, which are then sensedby the photodetector of the PET detector. Other radiotherapy systems,such as a proton therapy system, may also generate either scatteredX-rays or neutrons. Scattered radiation from proton sources may alsocause excitation of scintillation crystals. Afterglow of the PETdetectors 126 caused by scattered radiation (and/or other radiationsources) may accumulate over time and cause the detectors to saturate or“blank” for a period of time, rendering them incapable of detectingpositron emission event data during that blanking interval. FIG. 1Bdepicts an example of an output trace 110 from a PET detector 100, wherea linac pulse 101 was applied at t_(pulse). A linac pulse may have apulse width from about 1 μs to about 10 μs (e.g., from about 3 μs toabout 5 μs, about 3 μs about 5 μs about 8 μs etc.), with an inter-pulseinterval P_(interval) from about 2 ms to about 20 ms (e.g., from about 4ms to about 10 ms, from about 5 ms to about 15 ms, about 4 ms, about 10ms, etc.) and/or a pulse frequency from about 100 Hz to about 250 Hz.Scattered X-rays/radiation 102 from the linac pulse may irradiate thePET detectors 100, generating afterglow photons 104 in the scintillatingmaterial 103, which are then detected by photodetector 105. As seen inthe output trace 110, the afterglow photons cause a substantial shortterm artifact over a time period of from immediately after toapproximately 50 μs or more, e.g., 100 μs, during which time the PETdetector ability respond to positron emission events is reduced ordegraded. (e.g., PET detector saturation or blanking may be the resultof photodetector saturation, and/or the scintillator reaching itsmaximal photon output, and/or electrical and/or magnetic interferencefrom the linac, etc.). This time period may be referred to as theblanking interval 112, and is a short-term effect of detector afterglow.The blanking interval 112 may last from the beginning of the pulse toabout 50 μs (or more), depending on, for example, the duration andenergy of the linac pulse. After the initial blanking interval 112, thescintillating material of the PET detector may continue to scintillatesuch that afterglow photons continue to be generated, though possibly ata lower rate than during the blanking interval 112. These afterglowphotons may be generated by, for example, continued excitation and/orincreased energy levels of the scintillating material of the PETdetectors. The continued incidence of these afterglow photons on thephotodetector 105 may result in a greater level of noise 114 in theoutput trace 110 after the linac pulse was applied than before the pulsewas applied. This increased level of noise 114 may take about 1-5 hoursto decay to pre-linac pulse levels, and may be a long-term effect ofafterglow. In scenarios where there are high levels of scatteredradiation, afterglow photons may saturate the photodetector (e.g.,silicon photomultipliers). Since more than one linac pulse is emittedduring a treatment session (e.g., with about 2 ms to about 10 ms betweeneach pulse), afterglow noise of later pulses may cumulatively add to theafterglow noise of previous pulses, which may result in an increasinglynoisier signal on the output trace 110 of the PET detector. This maydisrupt the ability of the PET detectors to acquire accurate and precisepositron emission data throughout the duration of one or more treatmentsessions. In particular, the ability of PET detectors to detect a pairof coincident positron annihilation photons with sufficient precisionfor time-of-flight analysis may be compromised due to either the shortterm or long term afterglow effect.

Another way that the afterglow effect may disrupt the ability of the PETdetectors to acquire accurate and precise positron emission datathroughout the duration of one or more treatment sessions is from thedegradation of the energy resolution of the photodetector. As describedabove, photodetectors may saturate from afterglow photons. Aphotodetector, such as a silicon photomultiplier, may comprise hundredsto thousands of discrete Geiger avalanche photodiodes (which may bereferred to as micro-pixels). An optical photon that interacts with anindividual Geiger avalanche photodiode or micro-pixel may cause themicro-pixel to discharge. After discharging, the micro-pixel requiressome finite amount of time to recover. This finite amount of time may befrom about 10 ns to about 100 ns. If there is significant afterglow(e.g., as determined from an elevated bias current that exceeds athreshold), the total number of discrete micro-pixels available for thedetection of positron emission data may be reduced because they arefiring from afterglow photons, and cannot detect the scintillationsignal resulting from positron annihilation photons. As thephotodetector saturates from afterglow, its effective or cumulative gainis reduced. That is, the signal output from a photodetector affected byafterglow for a particular scintillation event is reduced as compared tothe signal output from a photodetector under normal (i.e.,non-afterglow) conditions. If the gain of the photodetector is reduced,then the quantitative accuracy of measuring the total energy of theincoming photon (e.g., scintillation event) may be degraded, which mayhinder the ability to reject scattered photons. While the sensitivity ofa PET detector may not be degraded by the afterglow effect, theafterglow effect may reduce the quantitative accuracy of the energy andtiming resolution of each scintillation event.

Afterglow may also cause photodetectors to detect or register positronannihilation photons (i.e., 511 keV photons) at a lower energy level;that is, instead of the photopeak of 511 keV photons being located atthe 511 keV level on the energy-spectrum, the photopeak of the 511 keVphotons are located at energy levels lower than 511 keV. Sincecoincidence detection controllers or processors are configured to detectpositron annihilation events based on 511 keV photons (e.g., setting adetection window centered around the 511 keV level), shifting thephotopeak of the 511 keV photons to a lower energy level (e.g., outsideof the detection window) may cause the PET detection system controlleror processor to miss the detection of a positron annihilation event.

Methods

One method for acquiring positron emission data from PET detectors inthe presence of scattered radiation may comprise adjusting the gain ofthe PET detector photodetectors (e.g., photomultipliers) as afterglow ofthe detectors increases, as depicted in the flow diagram of FIG. 2A. Asdepicted there, method 220 may comprise setting 222 an initial gainvalue for the photodetectors of the PET detectors of a radiation therapysystem. This step may take place during the assembly and/ormanufacturing of the system, or may take place just prior to the startof a radiation therapy session. After the gain value has been set, themethod may comprise proceeding 224 with radiation therapy, which maycomprise injecting a patient with a PET tracer, and activating a linacto generate and fire radiation pulses to one or more targets regions.During the radiation therapy session, a system controller may monitorone or more parameters and/or characteristics of the linac and/or PETdetectors and/or any other detectors or sensors (e.g., current orvoltage sensors, temperature sensors, radiation sensors, etc.). Thecontroller may determine 226 whether one or more of thosecharacteristics meet criteria for adjusting the gain value for thephotodetectors of the PET detectors. If one or more criteria have beenmet for adjusting the gain value, then the controller may adjust 228 thegain value of the PET detectors, e.g., by adjusting the bias voltage ofthe photodetectors and/or by adjusting a gain factor used in dataacquisition or analysis (e.g., adjusting an acquisition or analysissoftware gain factor) by a processor of the controller. For example, oneor more system parameters exceeding predetermined thresholds mayindicate that PET detector afterglow has increased to a certain level,and increasing the gain value of the photodetectors and/or the dataacquisition gain factor may help to reduce the false detection ofcoincident high-energy photons. At increased levels of afterglow, theremay be more photons generated by the scintillating material. Theseafterglow photons may cause the PET detector photodetectors to registerthe detection of 511 keV photons at lower energy levels. That is, theoutput from the PET detectors may indicate that photons at an energylevel lower than 511 keV were detected, when in fact, 511 keV photonswere detected, but the magnitude/energy of the PET detectorphotodetector output is reduced due to afterglow. Increasing the gainvalue of the photodetectors of the PET detectors (e.g., by increasingthe bias voltage to the photodetectors) may help to increase the PETdetector photodetector output so that it accurately reflects thedetection of 511 keV photons, which may help to improve the rate ofdetection of true coincident high-energy photons. Alternatively oradditionally, a gain factor may be used by the system processor in dataacquisition that compensates for the reduced PET detector output. Forexample, the system processor may multiply and/or shift the output of aPET detector by a gain factor whose value depends on the afterglowlevel. In some variations, the method depicted in FIG. 2A, along withthe methods depicted in FIGS. 3-7 may be implemented in machine-readableinstruction sets that may be stored in the memory of a controller incommunication with the PET detectors. Data from the radiation therapysystem, such as from various sensors, the PET detectors, the linac, etc.may be transmitted to the controller, which may perform computations(e.g., analysis) based on those measurements and/or may store theresults of those computations and/or system data in one or morecontroller memories. Command signals generated by the controller may betransmitted to the components of the radiation therapy system (e.g., thePET detectors and/or linac) to control the operation of those components(e.g., adjusting the gain value of the photodetectors of the PETdetectors).

One method for acquiring positron emission data from PET detectors inthe presence of scattered radiation may comprise adjusting thecoincidence trigger threshold of the PET detectors as afterglow of thedetectors increases, as depicted in the flow diagram of FIG. 2B. Asdepicted there, method 200 may comprise setting 202 an initialcoincidence trigger threshold for the PET detectors of a radiationtherapy system. This step may take place during the assembly and/ormanufacturing of the system, or may take place just prior to the startof a radiation therapy session. After the coincidence trigger thresholdhas been set, the method may comprise proceeding 204 with radiationtherapy, which may comprise injecting a patient with a PET tracer, andactivating a linac to generate and fire radiation pulses to one or moretargets regions. During the radiation therapy session, a systemcontroller may monitor one or more parameters and/or characteristics ofthe linac and/or PET detectors and/or any other detectors or sensors(e.g., current or voltage sensors, temperature sensors, radiationsensors, etc.). The controller may determine 206 whether one or more ofthose characteristics meet criteria for adjusting the coincidencetrigger threshold for the PET detectors. If one or more criteria havebeen met for adjusting the coincidence trigger threshold, then thecontroller may adjust 208 the coincidence trigger threshold of the PETdetectors. For example, one or more system parameters exceedingpredetermined thresholds may indicate that PET detector afterglow hasincreased to a certain level, and increasing the coincidence triggerthreshold may help to reduce the false detection of coincidenthigh-energy photons. That is, at increased levels of afterglow, theremay be more photons generated by the scintillating material. Theseafterglow photons may degrade or reduce the ability of the PET detectorsto detect coincident high-energy photons. Increasing the coincidencetrigger threshold of the PET detectors may help to disregard afterglowphotons, and help to improve the rate of detection of true coincidenthigh-energy photons. In some variations, the method depicted in FIG. 2B,along with the methods depicted in FIGS. 3-7 may be implemented inmachine-readable instruction sets that may be stored in the memory of acontroller in communication with the PET detectors. Data from theradiation therapy system, such as from various sensors, the PETdetectors, the linac, etc. may be transmitted to the controller, whichmay perform computations based on those measurements and/or may storethe results of those computations and/or system data in one or morecontroller memories. Command signals generated by the controller may betransmitted to the components of the radiation therapy system (e.g., thePET detectors and/or linac) to control the operation of those components(e.g., adjusting the coincidence trigger threshold of the PETdetectors).

The criteria for PET detector photodetector gain adjustment (e.g.,adjusting the gain value of the PET detector photodetectors and/or gainfactor used in positron emission data acquisition) and/or coincidencethreshold adjustment may be measured over an entire array of PETdetectors, and/or a PET detector module (i.e., having a subarray of PETdetectors), and/or a single PET detector. For example, in a radiationtherapy system with two PET detector arrays, each PET detector arraycomprising a plurality of PET detector modules (e.g., 32 PET detectormodules), each PET detector module comprising a subarray of PETdetectors (e.g., a 6×12 subarray of PET detectors), and where each PETdetector has its own photodetector, the criteria (and/or temperatures,bias currents, noise levels, coincidence timing distributions,photopeaks, dark count rates, etc.) may be measured over an entire PETdetector array, and/or over individual PET detector modules, and/or overindividual PET detectors. Similarly, the gain and/or coincidence triggerthreshold may be adjusted for an entire PET detector array, and/orindividual PET detector modules, and/or individual PET detectors. Forexample, all of the PET detectors in a PET detector module may have thesame photodetector gain value (i.e., bias voltage applied to the moduleis applied to all of the PET detector photodetectors), and a biascurrent measurement may be the cumulative bias currents of all of thePET detectors in the module. The bias current, bias voltage, and/or gainfactor for each PET detector module may be different from each other.That is, differing levels of afterglow correction may be applied todifferent PET detector modules. For example, in a radiation therapysystem with two PET detector arrays with 32 PET detector modules each,the afterglow effect may be corrected for each of the 64 PET detectormodules by measuring 64 bias currents of the 64 PET detector modules(and/or temperatures, noise levels, coincidence timing distributions,photopeaks, dark count rates, etc.) and then applying the afterglowcorrection to the 64 PET detector modules individually (e.g., applying64 potentially different gain and/or coincident threshold adjustments).Alternatively or additionally, the bias current (and/or temperatures,noise levels, coincidence timing distributions, photopeaks, dark countrates, etc.) may be measured for individual PET detector photodetectorsand/or over an entire PET detector array having multiple PET detectormodules. While the description and variations described below may referto measuring the bias current (and/or temperature, noise level,coincidence timing distribution, photopeak, dark count rate, etc.) for asingle PET detector and/or photodetector (or for a plurality of PETdetectors and/or photodetectors) and adjusting the gain and/or gainfactor and/or coincidence threshold for that single PET detector and/orphotodetector (or plurality of PET detectors and/or photodetectors,respectively), it should be understood that the description also appliesto measuring a plurality of bias currents (and/or temperatures, noiselevels, coincidence timing distributions, photopeaks, dark count rates,etc.) for a plurality of PET detectors and/or photodetectors (or for anindividual PET detector and/or photodetector), and adjusting the gainand/or gain factor and/or coincidence threshold for that plurality ofPET detectors and/or photodetectors (or for an individual PET detectorand/or photodetector, respectively).

One variation of a method for acquiring positron emission data in thepresence of scattered or stray radiation is depicted in FIG. 3. Method300 may comprise generating 302 a calibration table between detectornoise levels and coincidence trigger thresholds of PET detectors. Onemethod of generating a calibration table may comprise creatingenvironments that give rise to varying degrees or levels of noise on thePET detectors, providing a positron emission source (e.g., apositron-emitting seed) that emits positrons at a known rate, andadjusting the coincidence trigger threshold of the PET detectors at eachnoise level until the PET detector output corresponds to apre-determined time resolution quality or metric. The time resolutionquality or metric may be determined during the manufacture and/orcalibration of the radiation therapy system. The time resolution qualitymay be measured using a calibration source and may analyze the timespectrum of coincidence detected photons. For example, apositron-emitting point source may have a time-spectrum that follows aGaussian distribution where the mean is related to the spatial offset ofthe point source between PET detectors and the variance is related tothe quality of the time-resolving capability. One method for quantifyingthe time resolution quality may comprise calculating thefull-width-at-half maximum (FWHM) of this time spectrum. Method 300 mayalso comprise measuring 304 the noise level of the PET detectors duringa treatment session and comparing 306 the measured noise levels with thenoise levels in the calibration table to identify the coincidencetrigger threshold that corresponds with the measured noise level. Thecoincidence trigger threshold may be adjusted 308 based on changes inthe measured noise levels. For example, the coincidence triggerthreshold may be increased as the noise level on the PET detectorsincreases. Alternatively or additionally, the method 300 may be used toadjust the gain value of PET detector photodetectors and/or gain factorused in positron emission data acquisition (e.g., a gain factor used tomultiply and/or shift the output(s) of a PET detector(s)). For example,a variation of method 300 may comprise generating a calibration tablebetween detector noise levels and gain values and/or gain factors,measuring the noise level of the PET detectors during a treatmentsession and comparing the measured noise levels with the noise levels inthe calibration table to identify the gain values and/or gain factorsthat correspond with the measured noise level. The gain value and/orgain factor may be adjusted based on changes in the measured noiselevels.

Scattered X-rays may interfere with the ability of PET detectors toprecisely measure the arrival time of high-energy photons. In theabsence of scattered X-rays, the timing precision of PET detectors maybe characterized by a coincidence timing distribution having a range oftiming errors. The coincidence timing distribution may be measured, forexample, by using a point calibration source, as described above. Thetime difference from thousands or millions of coincidence events may beanalyzed and the coincidence timing distribution may be binned and/orhistogrammed to generate a timing distribution. Thefull-width-at-half-maximum (FWHM) of the timing distribution may be usedto characterized the timing resolution of a PET detector or entire PETsystem. As the levels of scattered radiation increase, the coincidencetiming distribution may change such that the range of timing errorsincreases. For example, without the interference of X-rays, PETdetectors may have a coincidence timing distribution such that the rangeof timing of errors is 300 ps FWHM, but in the presence of scatteredX-rays, the coincidence timing distribution may change such that therange of timing errors is 550 ps FWHM. One method of acquiring positronemission data in the presence of scattered radiation based oncoincidence timing distributions is depicted in FIG. 4. Method 400 maycomprise measuring 402 the coincidence timing distribution of the PETdetectors before the linac is activated (e.g., before a treatmentsession, and/or during manufacture and/or a calibration session),measuring 404 the coincidence timing distribution of the PET detectorsduring a time period when the linac has been activated (e.g., during atreatment session), and comparing 406 the coincidence timingdistribution measured in step 404 with the coincidence timingdistribution measured in step 402. If the timing distribution variesmore than about 10% from the previously measured timing distribution,then the coincidence trigger threshold of the PET detectors may beadjusted 408. One method of changing the coincidence trigger thresholdis by sending a command to a readout circuit (e.g., an ASIC) to increasethe voltage of a timing comparator. In another method, the coincidencetrigger threshold may be a predetermined number of optical photons thatare counted on a photodetector. In this method, the coincidence triggerthreshold may be adjusted by changing (e.g., increasing or decreasing)the number of photons that need to be detected to signal a coincidenceevent. Alternatively or additionally, the method 400 may be used toadjust the gain value of PET detector photodetectors and/or gain factorused in positron emission data acquisition (e.g., a gain factor used tomultiply and/or shift the output(s) of a PET detector(s)). For example,a variation of method 400 may comprise measuring the coincidence timingdistribution of the PET detectors before the linac is activated (e.g.,before a treatment session, and/or during manufacture and/or acalibration session), measuring the coincidence timing distribution ofthe PET detectors during a time period when the linac has been activated(e.g., during a treatment session), and comparing the coincidence timingdistribution measured during treatment with the coincidence timingdistribution measured before treatment. If the timing distributionvaries more than about 10% from the previously measured timingdistribution, then the gain value and/or gain factor may be adjustedbased on changes in the timing distribution.

Afterglow of PET detectors may cause the dark count rate of thephotodetector to increase, which may interfere with precise detection ofpositron emission events. Another variation of a method for acquiringpositron emission data in the presence of scattered radiation isdepicted in FIG. 5. In this method, the coincidence trigger thresholdmay be adjusted based on changes in the dark count rate of the PETphotodetector. Method 500 may comprise measuring 502 the dark count rateof the PET detectors before activation of the linac (e.g., before atreatment session, and/or during manufacture and/or a calibrationsession), measuring 504 the dark count rate of the PET detectors duringa time period when the linac has been activated (e.g., during atreatment session), and comparing 506 the dark count rate measured instep 504 with the dark count rate measured in 502. In some variations,the dark count rate may be measured by measuring the bias current of thephotodetector, and the comparison in step 506 may be between thecalculated dark count rate based on the bias current and/or the biascurrent measurement itself. Alternatively or additionally, the darkcount rate may be measured by counting low-photon triggers (i.e.,measuring the number of low energy photon triggers). The dark count ratemay be measured over an entire PET detector array or module/subarray,and/or may be measured on a per photodetector basis. If the dark countrate measured in steps 504 and 506 deviates more than about 2 Mcps toabout 10 Mcps (e.g., about 3 Mcps), the coincidence trigger threshold ofthe PET detector(s) may be adjusted 508. For example, the coincidencetrigger threshold of the PET detectors may be increased if the darkcount rate increases or exceeds a threshold (e.g., exceeds about 2 Mcps,exceeds about 3 Mcps, and/or exceeds about 10 Mcps). Steps 504-508 maybe repeated throughout the treatment session and/or while the linac isin use. Alternatively or additionally, the method 500 may be used toadjust the gain value of PET detector photodetectors and/or gain factorused in positron emission data acquisition (e.g., a gain factor used tomultiply and/or shift the output(s) of a PET detector(s)). For example,a variation of method 500 may comprise measuring the dark count rate ofthe PET detectors before the linac is activated (e.g., before atreatment session, and/or during manufacture and/or a calibrationsession), measuring the dark count rate of the PET detectors during atime period when the linac has been activated (e.g., during a treatmentsession), and comparing the dark count rate measured during treatmentwith the dark count rate measured before treatment. If the dark countrate measured during treatment deviates by more than about 2 Mcps toabout 10 Mcps (e.g., about 3 Mcps) from the dark count rate measuredbefore treatment, then the gain value and/or gain factor may be adjustedbased on changes in the timing distribution.

The effect of PET detector afterglow may also be measured in the biascurrent of the photodetector. Changes in the bias current may indicatedegradation in the ability of the PET detectors to acquire positronemission data, and adjusting the coincidence trigger threshold (e.g.,increasing the coincidence trigger threshold as the afterglow effectsincrease) may help improve the precision of the emission dataacquisition. One variation of a method for acquiring positron emissiondata in the presence of scattered radiation is depicted in FIG. 6A. Inthis method, the coincidence trigger threshold may be adjusted based onchanges in the bias current of the photodetector. The bias current maybe measured by monitoring the voltage supply (such as a high-voltagesupply) for the photodetector. Method 600 may comprise measuring 602 thebias current of the photodetector before activation of the linac (e.g.,before a treatment session, and/or during manufacture and/or acalibration session), measuring 604 the bias current of thephotodetector during a time period when the linac has been activated(e.g., during a treatment session), and comparing 606 the bias currentmeasured in step 604 with the bias current measured in 602. If the biascurrent measured in steps 604 and 606 deviate more than about 0.1 mA toabout 5 mA, the coincidence trigger threshold of the PET detectors maybe adjusted 608. Alternatively or additionally, the bias current may bemeasured over an entire PET detector array or module/subarray, and/ormay be measured on a per photodetector basis. Steps 604-608 may berepeated throughout the treatment session and/or while the linac is inuse.

Alternatively or additionally, the coincidence trigger threshold of thePET detectors and/or gain value of the PET detector photodetectorsand/or gain factor used in positron emission data acquisition may beadjusted based on temperature and/or radiation measurements of the areasat or around the linac (or any therapeutic radiation source) and/or PETdetector arrays. For example, a radiation therapy system may compriseone or more temperature sensors, which may be located at or near the PETdetector arrays and/or at or near the linac. Temperature data from thesesensors may be transmitted to the controller, and if the temperature atthe linac and/or the PET detector arrays exceeds one or more thresholds,the coincidence trigger threshold of the PET detectors may be adjusted.Similarly, one or more dosimeters (e.g., MOSFET dosimeter,thermoluminescent dosimeter, and the like) may be located at or near thePET detector arrays and/or at or near the linac. Radiation data fromthese dosimeters may be transmitted to the controller, and if theradiation levels at the linac and/or the PET detector exceed one or morethresholds, the coincidence trigger threshold of the PET detectors maybe adjusted. Some methods may also adjust the coincidence triggerthreshold and/or gain value of the PET detector photodetectors and/orgain factor used in positron emission data acquisition (e.g., a gainfactor used to multiply and/or shift the output(s) of a PET detector(s))based on the radiation output of the linac. For example, a radiationtherapy system may comprise a dose chamber or ionization chamberdisposed in the beam path of the linac. The ionization chamber maytransmit the amount of radiation emitted by the linac to the controller,which may adjust the coincidence trigger threshold of PET detectorsand/or gain value of the PET detector photodetectors and/or gain factorused in positron emission data acquisition based on the radiation outputof the linac. For example, a table that maps various radiation outputthresholds to various coincidence trigger thresholds and/or gain valueof the PET detector photodetectors and/or gain factor used in positronemission data acquisition may be stored in controller memory, and thecontroller may compare real-time ionization chamber measurements withthe thresholds in the table to determine whether to adjust thecoincidence trigger thresholds and/or gain value of the PET detectorphotodetectors and/or gain factor used in positron emission dataacquisition. The thresholds may be based on cumulative radiation outputstarting from the first pulse emitted by the linac until the currenttime point, and/or may be based on the radiation output over apredetermined interval of time (e.g., a pulse rate during a treatmentsession). For example, radiation output levels for a linac greater than0.1 Gy/min into a human torso may generate a sufficient level ofscattered radiation that may leads to afterglow in a PET detector.

In some variations, a table that maps linac pulse counts to variouscoincidence trigger thresholds and/or gain value of the PET detectorphotodetectors and/or gain factor used in positron emission dataacquisition (e.g., a gain factor used to multiply and/or shift theoutput(s) of a PET detector(s)) may be stored in controller memory. Thenumber of radiation pulses emitted by the linac may be used by thecontroller to adjust the coincidence trigger threshold of the PETdetectors. For example, the controller may adjust the coincidencetrigger threshold of the PET detectors after a first number of pulseshave been emitted by the linac, e.g., 10,000 pulses. The controller mayadjust the coincidence trigger threshold and/or gain value of the PETdetector photodetectors and/or gain factor used in positron emissiondata acquisition again when the linac has emitted an additional numberof pulses, e.g., another 10,000 pulses, bringing the cumulative pulsecount to 20,000. The number of pulses emitted by the linac (i.e.,threshold number of radiation pulses) before adjusting the coincidencetrigger threshold and/or gain value of the PET detector photodetectorsand/or gain factor used in positron emission data acquisition may beabout 1,000, about 2,000, about 4,000, about 7,500, or about 12,000pulses, etc., depending on level of scattered or stray radiation presentin a particular treatment system. That is, for systems with elevatedlevels of scattered or background radiation, the number of linac pulsesbefore adjusting the coincidence trigger threshold and/or gain value ofthe PET detector photodetectors and/or gain factor used in positronemission data acquisition may be lower than for systems with lowerlevels of scattered or background radiation. In some variations, thetable may map linac pulse rates or pulse schedules (i.e., number ofpulses over a particular interval of time, and/or a timing schedule ofpulses) to PET detector coincidence trigger thresholds and/or gainvalues of the PET detector photodetectors and/or gain factor used inpositron emission data acquisition. One or more of these parameters maybe used alone and/or in combination with one or more of the methodsdescribed herein to determine when to adjust PET detector coincidencetrigger thresholds and/or gain values of the PET detector photodetectorsand/or gain factor used in positron emission data acquisition, and/orhow much to adjust the coincidence trigger thresholds (e.g., increase ordecrease by a specific value, etc.). As an example, the initialcoincidence trigger threshold for PET detectors at the beginning of atreatment session may be about 2 photon-triggers. A photon-trigger maybe the voltage, charge or count that represents a detected photon. Forexample, a 2 photon-trigger means that the timing discriminator of a PETdetector fires when it detects the arrival of two or more photons. After10,000 radiation pulses have been emitted, the coincidence triggerthreshold may be increased to about 5 photon-triggers. After another10,000 radiation pulses have been emitted (that is, 20,000 radiationpulses cumulatively), the coincidence trigger threshold may be increasedto about 6 photon-triggers. The threshold number of radiation pulsesbefore changing the coincidence trigger threshold, as well as thecoincidence trigger threshold change increments may vary from thisexample, as may be desirable.

FIG. 6B depicts one variation of a method where the gain value of PETdetector photodetectors may be adjusted based on changes in the biascurrent of the photodetector, which may compensate for the saturation ofthe photodetector (e.g., silicon-photomultiplier) due to afterglow. Thebias current may be measured by monitoring the voltage supply (such as ahigh-voltage supply) for the photodetector. Method 620 may comprisemeasuring 622 the bias current of the photodetector before activation ofthe linac (e.g., before a treatment session, and/or during manufactureand/or a calibration session), measuring 624 the bias current of thephotodetector during a time period when the linac has been activated(e.g., during a treatment session), and comparing 626 the bias currentmeasured in step 624 with the bias current measured in 622. If the biascurrent measured in steps 624 and 626 deviate more than about 0.1 mA toabout 5 mA, a gain value of the PET detector photodetectors and/or gainfactor used in positron emission data acquisition may be adjusted 628.The bias current may be measured over an entire PET detector array ormodule/subarray, and/or may be measured on a per photodetector basis.Steps 624-628 may be repeated throughout the treatment session and/orwhile the linac is in use. In some variations, the gain value of thephotodetector(s) may be adjusted by adjusting the bias voltage of thephotodetector and/or a gain factor used in positron emission dataacquisition (e.g., adjusting an acquisition or analysis software gainfactor) by a processor of the controller. For example, a gain factor maybe used by the system processor in data acquisition to help compensatefor the reduced PET detector output due to the afterglow effect. Forexample, the system processor may multiply and/or shift the output of aPET detector by a gain factor whose value depends on the afterglow level(as indicated by, for example, the measured bias current). The gainfactor may be calculated by, for example, measuring PET detector outputvalues corresponding to 511 keV photons at different levels of afterglow(corresponding to different values of measured bias current), taking adifference between the measured PET detector output values and thenominal output value corresponding to the detection of 511 keV photons(i.e., in the absence of afterglow), and calculating a ratio of thedifference to the nominal output value. Alternatively or additionally,the gain factor may be calculated by measuring a photopeak shift of 511keV photons at different levels of afterglow (corresponding to differentvalues of measured bias current), taking a difference between theshifted photopeak(s) and the nominal photopeak at 511 keV (i.e., in theabsence of afterglow), and calculating a ratio of the difference to eachof the shifted photopeaks. A calibration table or plot may be generatedthat maps the measured bias current to gain factor values. Changing thegain value and/or gain factor used in positron emission data acquisitionmay adjust the energy and timestamp of a positron annihilation eventrecorded by the PET detectors to correct for saturation of the detectorcaused by afterglow. In some variations, a method for adjusting the gainvalue and/or gain factor may comprise generating a calibration tablebetween bias current values and gain values (and/or gain factors),measuring the bias current of the PET detector photodetectors during atreatment session, comparing the measured bias current with the biascurrent values in the calibration table to identify the gain valuesand/or gain factors that correspond with the measured bias current, andthen adjusting the gain value and/or gain factor according to thecalibration table.

FIG. 13A depicts experimental data plots of the changes in bias currentand temperature as a function of linac beam over time, while keeping thebias voltage constant. As depicted there, it can be seen that after thebeam is turned on at time point 0, the bias current increases from a lowlevel of approximately 0.2 mA to approximately 4.2 mA over about onehour. When the beam is turned off at time point 69 (i.e., 69 minutesafter the beam was turned on), the bias current drifts downward back toits baseline value over about 100 minutes. The temperature also driftsupward as the PET detector heats up because it has to dissipate morepower in the photodetector. The heat accumulated in the PET detector maybe the bias-current multiplied by the bias voltage. In this experiment,the heat generated in the detector starts at approximately 11 mW ataround time point 0 (0.2 mA*55V) but increases significantly toapproximately 231 mW at the peak around time point 69 (4.2 mA*55V). Thechange in temperature also affect the gain value of the photodetector,which may be corrected or compensated by adjusting the gain factor usedin positron emission data acquisition, as described herein.

A calibration table may be generated by measuring positron emission dataof a calibration positron emitting point source (e.g., Na-22) using aPET detector, and tracking (i.e., quantifying) how that measurementchanges at differing levels of afterglow. FIG. 13B depicts a calibrationplot (which may be presented as a calibration table) that has beengenerated by measuring the bias current and photopeak location over aPET detector module having a 6×12 subarray of PET detectors. Thephotopeak is the location of the 511 keV peak of the photodetector alongthe energy-spectrum as detected by a PET detector photodetector. Photonsemitted by a positron annihilation event have the same energy, and sotherefore, every valid event has the same energy value. The “photopeak”may be the 511 keV peak on an energy-spectrum histogram plot of all thedetected events. As can be seen in FIG. 13B, as the photodetectors ofthe PET detector is subject to greater afterglow, the bias currentincreases and the photopeak of 511 keV photons is detected as iflower-energy photons were detected (e.g., drifts linearly downward from511 keV (nominal) to 400 keV). This downward shift of the 511 keVphotopeak may degrade the performance of the photodetector, and hinderits ability to recognize or report 511 keV photons resulting from apositron annihilation event. FIG. 13C depicts a plot of the energyresolution of a PET detector over time (each data-series intervalrepresents a 10 minute increment, where the linac beam was turned on atdata-series value 1 and turned off at data-series value 7), where thegain of the PET detector photodetector has been adjusted to compensatefor the downward shift of the 511 keV photopeak. As depicted there, withthe gain adjustment or correction, the energy resolution remains stableover time. As described previously, photodetector gain may be adjustedor changed by adjusting the bias voltage to the detector. Alternativelyor additionally, a gain factor that may be used by the system processorin data acquisition may be adjusted so that the system processormultiplies and/or shifts the output of a PET detector by the gainfactor. In one variation, the gain factor for each bias current valuemay be the slope of the curve or line at that bias current value in aphotopeak location to bias current plot. In the example of FIG. 13B, thegain factor may be the slope of the line (fitted to the data points inthe plot), and the gain factor may be multiplied with the PET detectormodule output to identify positron annihilation events. A calibrationtable may be generated based on the plot of FIG. 13B that maps measuredbias current levels to the photopeak location on the energy-spectrum.FIG. 13C shows the result of applying gain correction under differentafterglow scenarios. As depicted there, adjusting the gain factor indifferent afterglow conditions/levels helps to keep the energyresolution or quality of measuring the energy of the incoming photonrelatively constant. Monitoring the location of the photopeak on theenergy-spectrum (e.g., during a treatment session) may provide anindication as to whether the gain has been appropriately adjusted tocompensate or correct for the effects of afterglow (e.g., gain is nottoo high or too low). For example, alternatively or additionally to acalibration table that maps bias current levels to gain values (e.g.,bias voltage levels, software gain factor), a calibration table may begenerated that maps photopeak locations (e.g., 511 keV photopeaklocations) to gain values (and/or gain factors used in positron emissiondata acquisition) so that during treatment, the gain value and/or gainfactor may be adjusted according to photopeak locations.

Another metric that may be used to determine whether the photodetectorgain adjustment appropriately corrects for the afterglow effect is thetime resolution of the photodetector. The time resolution of aphotodetector, which represents the smallest time interval between twophoton detection events that may be distinguishable by the photodetectoras two separate events, may shift due to the afterglow effect. FIG. 13Ddepicts the shift of a time resolution centroid over time (eachdata-series interval represents a 10 minute increment, where the linacbeam was turned on at data-series value 1, turned off at data-seriesvalue 7, and the recovery period is from data-series values 8-16) as aPET detector photodetector is subject to scattered linac X-rays thatcause afterglow. The shift value of the timing centroid may be used, insome variations, to predict or estimate afterglow levels. Some methodsfor adjusting the gain of photodetectors to correct for aftergloweffects may comprise calculating a calibration table that maps shifts inthe timing centroid shift as a function of bias current. Alternativelyor additionally, a calibration table may be generated that maps timingcentroid shifts to gain values (e.g., bias voltage levels, software gainfactor). During a treatment session, the timing resolution (e.g., timingcentroid) may be measured/monitored and may provide an indication as towhether the gain has been appropriately adjusted to compensate orcorrect for the effects of afterglow (e.g., gain is not too high or toolow). In some variations, the gain of PET detector photodetectors may beadjusted to correct for any timing centroid drifts.

As described previously, any of the methods for acquiring positronemission data from PET detectors in the presence of scattered radiationmay comprise measuring and monitoring one or more parameters and/orcharacteristics of the linac and/or PET detectors and/or any otherdetectors or sensors (e.g., current or voltage sensors, temperaturesensors, radiation sensors, etc.), and determining the afterglow levelor severity based on these one or more parameters. That is, parameterssuch as temperature, bias current, radiation emission levels and/orpulse count, etc. may act as surrogates that quantify the afterglowlevel or effect. Based upon these measurements, a treatment system maychange the gain to PET detector photodetectors by applying changes tothe bias voltage and/or corrections or changes to the gain value used indata acquisition or analysis (e.g., adjusting an acquisition or analysissoftware gain factor) by a processor of the controller.

Alternatively or additionally, some methods may comprise delaying theacquisition of PET data by the controller during a linac radiation pulseand for a specified time interval after the radiation pulse. Delaying orsuspending PET data acquisition and/or transmission during a linac pulseand for a specified time interval after the pulse may help to reduce oreliminate the storage and processing of positron emission data withafterglow noise and/or radiation pulse artifacts. The amount ofradiation artifacts from the linac pulse and/or afterglow effect may bethe greatest during the pulse and the time interval immediatelyfollowing the pulse, and processing positron emission data with elevatedlevels of noise or artifacts may result in incorrect or imprecisecoincidence detection. In some variations, the width of the radiationpulse from the linac may be about 5 μs or less and pulsed at a frequencyof about 100 to about 300 Hz. In this configuration, the duty cycle ofactual radiation beam on-time is from about 0.05% to about 0.15% (i.e.,radiation beam off-time is about 99.85% to about 99.95%). The PETdetectors and/or controller may delay and/or gate the acquisition of PETdata during the linac pulse (e.g., delaying and/or gating theacquisition of positron emission data during a 5 μs linac beam pulse)and/or a period of time after the linac pulse with little or no impacton PET sensitivity. For time-of-flight PET systems, reducing oreliminating relatively high-noise positron emission data from thetime-of-flight calculations may help to facilitate more precise locationcalculations, and/or may help to reduce margins of error.

The duration of the delay time interval may be determined at least inpart based on the amount of PET detector afterglow, which may bequalitatively and/or quantitatively determined based on one or more ofPET detector noise levels, detector timing distributions, dark countrate, bias current, temperature, ambient radiation levels, etc.,including any of the parameters described previously. For example, thedelay time interval may be from about 85 μs to about 500 μs, e.g., atleast about 100 μs, at least about 200 μs, etc. In some variations,delaying the acquisition of positron emission data by the controller maycomprise gating the reception of positron emission data by thecontroller such that positron emission data are not stored by thecontroller if the data was detected by the PET detectors during thespecified time interval after a linac pulse. Alternatively oradditionally, the transmission of positron emission data from the PETdetector to the controller may be delayed such that PET data detected bythe PET detectors during the specified time interval after a linac pulseis not transmitted. For example, data transmission from the PETdetectors to the controller may be suspended during the specified timeinterval after a linac pulse, and may resume after the specified timeinterval has elapsed. In some variations, delaying the acquisition ofpositron emission data by the controller may comprise reading outpositron emission data stored by the controller after the specified timeinterval after a linac pulse. For example, the positron emission datamay be acquired and stored in controller memory even during thespecified time interval after a linac pulse, however, the controllerdoes not read the positron emission data from the memory until the afterthe specified time interval has passed and the positron emission datastored in the controller memory reflects data acquired after thespecified time interval.

FIG. 7A depicts a flowchart representation of one variation of a methodfor gating positron emission data. As depicted there, method 700 maycomprise measuring 702 the effects of afterglow on the PET detectors.The amount of afterglow may be qualitatively and/or quantitativelydetermined based on one or more of the metrics described herein,including but not limited to PET detector noise levels, detector timingdistributions, dark count rate, bias current, temperature, ambientradiation levels, etc. The method 700 may comprise setting 704 a gatetrigger threshold based on the amount of afterglow. A gate triggerthreshold is the time interval during which positron emission data isnot transmitted from the PET detectors and/or positron emission data isnot stored by the controller. For example, if any of the parametersindicative of afterglow levels exceed a predetermined threshold (e.g., anoise level, timing, dark count rate, bias current level, temperaturelevel, radiation level), the gate trigger threshold may be increasedsuch that communication of positron emission data between the PETdetectors and the controller is suspended for a longer period of time.The gate trigger threshold may remain at an increased level until theafterglow returns to the PET detector pre-afterglow level (or withinabout 5% of the pre-afterglow level). The method 700 may furthercomprise gating 706 the positron emission data during and/or after thelinac pulse based on the gate trigger threshold. The method 700 may becontinuously executed by the controller during a treatment session,and/or may be executed at prescribed time intervals. In some variations,the method 700 may further comprise changing or updating the gatetrigger threshold from an initial level if the measured afterglow effectexceeds a selected threshold, and/or if the change in afterglow from aprevious measurement exceeds a selected threshold. For example, if theamount of detector afterglow remains relatively constant during atreatment session, the gate trigger threshold may not be updated.However, if there are substantial shifts in the amount of detectorafterglow (e.g., increases or decreases in detector afterglow), the gatetrigger threshold may be updated. The gating may be increased from about500 μs to up to about 10 ms. Alternatively, the gate trigger thresholdmay be held constant.

FIG. 7B depicts a schematic diagram of one variation of a logic circuit709 for gating PET data communication between the PET detectors and aprocessor of the controller. One or more PET detectors 710 (in one ormore PET detector arrays) may output data 711 to a comparator 712. Thecomparator 712 compares the timing characteristics of the positronemission data 711 with a coincidence trigger threshold 720 from aprocessor 716 of the controller. The coincidence trigger threshold 720may be determined when the system initiates (e.g., may initiate to apreset, default value) and/or may be updated as the system is used fortreatment, as described above. If positron emission data is within thecoincidence trigger threshold, then the positron emission data is outputto a gate 714. Gate 714 may be, for example, an “AND” logic gate.Positron emission data may be transferred to the processor 716 only if agate signal 718 from the processor to the gate 714 is active (e.g.,“high”). The timing of the gate signal 718 may be such that it isinactive (e.g., “low”) during a linac pulse and/or an interval of time(e.g., delay time) after the pulse, and is active (e.g., “high” afterthe interval of time following a linac pulse has elapsed and/or before alinac pulse. FIG. 7C depicts one example of a timing diagram 730 thatdepicts the timing of signals between the linac pulse 732, the inverseof the gate signal (which may be considered a “reset” signal), and acommon clock signal 736 shared between the linac, PET detectors and theprocessor. In this example, the linac pulse 732 may have a pulse widthof about 5 μs. In the variation depicted the timing diagram 730, theinverse of the gate signal 734 is “high” during a setup time periodt_(setup) of about 10 μs before the start (i.e., rising edge) of a linacpulse 733, during the linac pulse 733, and for a time interval (e.g.,delay time) after the end (i.e., falling edge) of the linac pulse 733.The linac pulse 733 may be from about 2 μs to about 15 μs, with a dutycycle of less than about 0.001. The total duration t_(gate) that theinverse of the gate signal is “high” may be from about 100 μs to about 3ms. While the inverse of the gate signal is “high”, positron emissiondata transfer/acquisition is suspended. Positron emission datatransfer/acquisition may resume when the inverse of the gate signal is“low”. The duration of time between gate signal pulses (i.e., interpulseperiod t_(interpulse)) may be as low as about 1 ms, but in somevariations, the interpulse period between linac pulses may be betweenabout 3 ms to about 10 ms. The time interval following the end of thelinac pulse 733 may be adjusted based on the degree of PET detectorafterglow, as described above.

Alternatively or additionally, the system may interleave positronemission data acquisition with radiotherapy delivery (e.g., linacactivation). In this method, the PET detectors first acquire positronemission data. In some variations, the positron emission data from thePET detectors may be used to generate an image. After the positronemission data has been acquired and stored (e.g., into a controllermemory), and/or after a PET image has been generated using the positronemission data, the PET detectors may be de-activated or disabled. Theradiation source (e.g., linac or proton source) may be activated afterthe PET detectors have been de-activated, and may emit radiation pulsesto the target (e.g., tumor region). In this interleaved mode, thepositron emission data acquisition and the radiotherapy beam emission donot overlap significantly in time. In some variations, the activation ofthe PET detectors and the radiation source may be on a 50/50 duty cycle.While this may allow for longer periods of positron emission dataacquisition, it may extend the overall length of the treatment session.

The PET system may interlock or cease positron emission data acquisitionafterglow levels exceed a predetermined threshold. In one variation, thebias current of the PET detector photodetectors may be measured and ifthe bias current exceeds a predetermined bias current interlockthreshold, the PET detector(s) may interlock and cease data acquisition.The system controller may continue to poll the bias current at regularintervals, optionally generating notifications to a clinician ortechnician that indicate the bias current levels and/or afterglowlevels. The PET detector(s) may resume data collection (i.e., clear theinterlock) when the bias current value is lower than aninterlock-release threshold. In some variations, the interlock-releasethreshold may be the same as the interlock threshold, while in othervariations, the interlock-release threshold may be less than (e.g.,lower than) the interlock threshold.

Oscillating Scatter Shield

Some variations of a radiotherapy system may comprise a movableradiation shield or filter that may be disposed over the PET detectorsduring a radiation pulse and moved away from the PET detectors after theradiation pulse. The radiation shield or filter may absorb scatteredradiation and/or deflect scattered radiation away from PET detectors,which may help to reduce the amount of PET detector afterglow. Aphysical shield or filter that eclipses the PET detectors during linacpulses may also help reduce or eliminate the back-projection informationassociated with the linac pulses. In some variations, the shield mayeclipse the PET detectors during the linac pulses and expose the PETdetectors before or after the linac pulses. Because of the inertiaassociated with a physical shield or filter, some of the PET detectorsmay be eclipsed for more time than the linac pulse.

For example, emission-guided radiation therapy systems may comprise aplurality of PET detectors and a linac that are mounted on a gantry thatis rotatable about a patient. Emission data acquired in real-time by thedetectors may be analyzed by a system controller. Based on this emissiondata, the system controller may rotate the gantry to direct radiationfrom the linac to the PET-avid tumor regions from various firing angles.In some variations, the linac and the PET detectors may be mounted on arotatable ring-shaped or circular gantry and the patient treatment areamay be located along the center of the circular gantry (e.g., along theaxis of rotation). A radiation therapy system may comprise a radiationfilter ring, which may comprise one or more radiation shields orfilters. In some variations, the radiation filter ring may comprise aclosed ring structure, while in other variations, the radiation filterring may comprise one or more ring segments (e.g., open ring, partialsegments or arcs of a ring, etc.). The radiation filter ring may besized to fit within an inner diameter of the first circular gantry. Theradiation filter ring may have the same axis of rotation as the firstcircular gantry, and may move independently from the circular gantry. Insome variations, the radiation filter ring may rotate while in othervariations, the radiation filter ring may oscillate with respect to thecircular gantry, where the radiation filter ring moved into the diameterof the gantry and out of the inner diameter of the gantry in a lateraldirection along the axis of rotation of the gantry. In some variations,radiation shields or filters may comprise one or more radiation-blockingor radiopaque components (e.g., panels comprising high-Z materials) thatmay be circumferentially located along the radiation filter ring. Theother portions of the radiation filter ring may beradiation-transmitting or radiotransparent (e.g., comprising low-Zmaterials). The radiation-blocking section(s) of the radiation filterring (i.e., the portions of the radiation filter ring where theradiation-blocking components are located) may have a size and shapethat corresponds to the size and shape of the PET sensor arrays. In afirst configuration (e.g., a radiation-blocking configuration), theradiation filter ring may be located such that the radiation-blockingsections are disposed over the PET detector arrays. In a secondconfiguration (e.g., a radiation-transparent configuration), theradiation filter ring may be located such that the radiotransparentsections (i.e., the portions of the radiation filter ring which do nothave radiation-blocking components) are disposed over the PET detectorarrays. A motion controller (that may be separate and/or independentfrom the motion controller for the first circular gantry) coupled to theradiation filter ring may rotate or oscillate the radiation filter ringto transition between the first and second configuration. The motioncontroller for the radiation filter ring may comprise an actuator,motor, and/or drive mechanism coupled to the filter ring that supplies amotive force sufficient for changing the position of the filter ringaccording to a specific time interval or schedule (as described furtherbelow). In some variations, the motion controller may comprise a springmechanism having one or more springs and an actuator system or mechanism(e.g., a pneumatic or hydraulic actuator, cam-based motor, slotted-linkmotor, electromagnetic actuators, etc.). The spring mechanism may assistthe actuator system or mechanism by providing an additional motive forceto help expedite filter ring motion and/or offset any energy loss in themotion system due to frictional and/or drag forces.

In some variations, a rotatable radiation filter ring may make one halfof a rotation for each linac pulse. Alternatively or additionally, anoscillating radiation filter ring may make half a cycle for each linacpulse. An oscillating radiation filter ring may be centered over the PETdetector arrays so that its velocity (e.g., the lateral speed at whichit moves across the PET detector arrays into and out of the innerdiameter of the gantry) is higher when the radiation-blocking sectionseclipse the PET detectors as compared to its velocity when theradiation-transmitting sections are disposed over the PET detectors. Thetime during which the PET detector arrays are obstructed or shieldedfrom scattered radiation (e.g., from a linac pulse) is relatively shortas compared to the time during which the PET detector arrays areunobstructed. That is, for a given linac pulse duty cycle, the PETdetectors are in PET data acquisition mode and as such, may beunobstructed by the radiation-blocking sections of the radiation filterring. For example, the motion controller may be synchronized with thelinac such that it moves the radiation filter ring into the firstconfiguration during a linac pulse (and optionally, for a time periodbefore and/or after the pulse) and it moves the radiation filter ringinto the second configuration after the pulse (e.g., during theinter-pulse interval). The duty cycle, pulse width, pulse frequency ofthe linac pulse may be communicated to the radiation filter ring motioncontroller such that the radiation filter ring is in the firstconfiguration during a linac pulse. In some variations, a motioncontroller may oscillate the radiation filter ring to cover severaltimes the PET detector array width in its path. The length of theoscillation displacement of the radiation filter ring (e.g., thecircumferential or arc length of swept by the radiation filter ring asit oscillates between the first configuration and the secondconfiguration) may be selected such that it is greater than the width ofa PET detector array. This may help to reduce the proportion of the PETdetectors that are obscured when the oscillating radiation filter ringis in the first configuration. For example, an oscillation displacementhaving a length that is nine times the width of the PET detector arraywidth eclipses only 5% of the available PET events.

FIGS. 8A-8B depict one variation of a radiation therapy system 800comprising PET detector arrays 802 a, 802 b and a rotatable radiationfilter ring 804 around a bore 803 of the system. The rotatable radiationfilter ring 804 comprises a first radiation-blocking section 806 a and asecond radiation-blocking section 806 b. The diameter of the rotatableradiation filter ring may be such that it that approximates the innerdiameter of a circular gantry upon which the PET detector arrays aremounted. In some variations, the diameter of the rotatable radiationfilter ring may be less than the inner diameter of the circular gantry.The PET detector arrays 802 a, 802 b may be located across from eachother (e.g., opposite each other, such that the center of the first PETdetector array 802 a is located about 180 degrees from the center of thesecond PET detector array 802 b). Similarly, to correspond with thearrangement of the PET detector arrays 802 a, 802 b, theradiation-blocking sections 806 a, 806 b may also be located across fromeach other. The length of the radiation-blocking section 806 a, 806 bmay be selected such that their circumferential length corresponds withthe circumferential length of the PET detector arrays 802 a, 802 b. FIG.8A depicts a first configuration of rotatable radiation filter ring 804where the radiation-blocking sections 806 a, 806 b (which may compriseone or more high-Z materials) are disposed or aligned over the PETdetector arrays 802 a, 802 b. For example, the rotatable radiationfilter ring may be in this first configuration during a linac pulse.FIG. 8B depicts a second configuration of the rotatable radiation filterring 804 where the radiation-blocking sections 806 a, 806 b are notdisposed over (i.e., are misaligned to) the PET detector arrays when thelinac is inactive (e.g., not emitting a pulse, or within an inter-pulseinterval). FIG. 8C is a schematic cross-sectional view of the system 800of FIGS. 8A-8B. FIG. 8C depicts the rotatable radiation filter ring inthe first configuration, where the radiation-blocking sections aredisposed over the PET detector arrays such that the PET detector arrays802 a, 802 b are shielded from radiation along a treatment plane 801through the bore 803.

FIGS. 9A-9B depict a schematic cross-sectional view of another variationof a radiation filter ring 900, where instead of rotating the radiationfilter ring within the inner diameter of the circular gantry (which maybe disposed around a bore 903), the radiation filter ring 900 laterallyoscillates into and out the inner diameter of the circular gantry. Theoscillating radiation filter ring 900 may comprise first and secondradiation-blocking sections 902 a, 902 b, which may have a size andshape that corresponds with the size and shape of the PET detectorarrays 904 a, 904 b. FIG. 9A depicts a first configuration of theoscillating radiation filter ring 900 where the oscillating radiationfilter ring is located within the inner diameter of the circular gantry(i.e., the radiation-blocking sections are disposed over the PETdetector arrays). As described above, the radiation-blocking sectionsmay comprise one or more panels of high-Z materials for reflectingand/or absorbing scattered radiation, and when disposed over the PETdetector arrays, may help to reduce PET detector afterglow by reducingor eliminating the incidence of scattered radiation on the PETdetectors. FIG. 9B depicts a second configuration of the oscillatingradiation filter ring, where the radiation-blocking sections 902 a, 902b are not disposed over (i.e., are not aligned to) the PET detectorarrays. The lateral movement of the oscillating radiation filter ring isindicated by the arrows 901. The oscillating radiation filter ring maybe in the first configuration during the linac pulse such that the PETdetector arrays 904 a, 904 b are shielded from radiation along atreatment plane 901 through the bore 903, and may be in the secondconfiguration when the linac is inactive (e.g., not emitting a pulse, orwithin an inter-pulse interval).

While the variations of rotatable or oscillating radiation filter ringsdescribed above are circular or ring-shaped, in other variations,radiation filters may be blocks of radiation-blocking or radiopaquematerials that are moved over the PET detectors during time intervalswhere high levels of scattered radiation are expected and moved awayfrom the PET detectors during time intervals where relatively low levelsof scattered radiation are expected. For example, radiation filters maybe mounted on arms, rails, etc. and/or may be coupled to actuators ormotors that move them over and away from the PET detectors.

Examples

FIG. 10A depicts one example of an experimental setup where the PETdetector afterglow was measured and characterized. Short term and longterm afterglow on the PET detectors were measured. An acrylic scatteringtarget 1000 was used to scatter radiation from a 6 MeV linac 1002. Thetarget 1000 was a 40 cm×40 cm×20 cm block of acrylic. The scatteredradiation incident on two PET detectors run in coincidence was measuredusing a PET detector 1004, to simulate the radiation scattering of apatient. The center of the target 1000 was 53 cm from the PET detector1004. The 6 MeV linac 1002 generated a 10 cm×10 cm field of X-rayradiation onto the acrylic phantom 1000 with a 3 us pulse width at 153Hz. The entire setup was enclosed by a plurality of lead blocks 1006,each having a length of about 20 cm. There was a first dose chamber 1008in the path of the linac beam 1010, a second dose chamber 1008 locatedbehind the PET detectors 1004, and a dose chamber 1008 located on thewall of the target 1000 furthest from the linac 1002. A film 1009 waslocated on the wall of the target 1000 closest to the linac 1002. FIG.10B is a schematic depiction of two single crystal PET detectors 1012(with LYSO scintillation crystal) and 1014 (with LFS scintillationcrystal) that were used with the setup of FIG. 10A. Both of the PETdetectors 1012, 1014 used a solid state photodetector or photomultiplier1016 (such as MPPC (SiPM) photodetectors), which were capable oftime-of-flight measurements, and were used to measure the effect ofscattered radiation on PET data acquisition.

The silicon photomultiplier (SiPM) used in the PET detector may becharacterized by its dark-count rate performance. The SiPM may besensitive to single photons, and dark-counts are thermionic noise eventsin the detector. The scintillation detector produces a short-termafterglow, and these generate optical photons that persist betweenscintillation pulses. These approximate the dark-counts from thepoint-of-view of the detection system. The dark-count rate (DCR) of thesensor was characterized, as a function of when it occurred relative tothe LINAC pulse (FIG. 10C). The DCR right before a linac pulse and theDCR 50 μs to 200 μs after the linac pulse were characterized. Based onthese analyses, delaying the acquisition of PET data by the controllerby about 200 μs helped to reduce the DCR to an acceptable level.

The initial DCR for the photo-sensor was 2M dark-counts per second. Thismatched the specification from the vendor for the device. FIG. 10Ddepicts the DCR as a function of total time exposure to the linac for aLYSO scintillation crystal. As the linac is pulsed, the minimum DCR wasimmediately before the linac pulse as expected (FIG. 10D, bottom trace).If the DCR is measured immediately after the linac pulse, it issignificantly higher. Therefore, there is a short-time constant decaythat contributes to afterglow.

FIG. 10E depicts DCR as a function of total time exposure to the linacfor a LYSO scintillation crystal for two different crystals: LFS andLYSO. The linac is run for 3600 seconds or 1 hour continuously. As thelinac is run over time, the afterglow tends to build up in thescintillation crystal. This after-glow cannot be distinguished fromdark-counts of the sensor. It significantly degrades the signal-to-noiseratio of the SiPM. Note, the sensor starts at a DCR of approximately 2Mdark-counts/seconds. After 1 hour of afterglow, it increases to wellover 10M dark-counts per second, or a factor of 5 worse.

FIG. 11 depicts another example of an experimental setup for measuringafterglow using two multi-crystal PET detectors.

FIG. 12 depicts a plot of a time-of-flight PET module at various timepoints after a linac pulse, where each curve depicts the time resolutionat different coincidence trigger thresholds. As shown, afterglow noisemay be mitigated by increasing the threshold of the PET detector modules(FIG. 11). The afterglow effect was mitigated by changing the triggerthreshold of the PET detector electronics. It may be possible to measurethe DCR of the system during therapy and adapt the threshold to helpimprove the timing performance of the system.

The radiation treatment systems described herein may comprise acontroller having a processor and one or more memories. A controller maycomprise one or more processors and one or more machine-readablememories in communication with the one or more processors. Thecontroller may be connected to a radiation therapy system and/or othersystems by wired or wireless communication channels. In some variations,the controller of a radiation treatment system may be located in thesame or different room as the patient. For example, the controller maybe coupled to a patient platform or disposed on a trolley or medicalcart adjacent to the patient and/or operator.

The controller may be implemented consistent with numerous generalpurpose or special purpose computing systems or configurations. Variousexemplary computing systems, environments, and/or configurations thatmay be suitable for use with the systems and devices disclosed hereinmay include, but are not limited to software or other components withinor embodied on personal computing devices, network appliances, serversor server computing devices such as routing/connectivity components,portable (e.g., hand-held) or laptop devices, multiprocessor systems,microprocessor-based systems, and distributed computing networks.

Examples of portable computing devices include smartphones, personaldigital assistants (PDAs), cell phones, tablet PCs, phablets (personalcomputing devices that are larger than a smartphone, but smaller than atablet), wearable computers taking the form of smartwatches, portablemusic devices, and the like.

In some embodiments, a processor may be any suitable processing deviceconfigured to run and/or execute a set of instructions or code and mayinclude one or more data processors, image processors, graphicsprocessing units, physics processing units, digital signal processors,and/or central processing units. The processor may be, for example, ageneral purpose processor, Field Programmable Gate Array (FPGA), anApplication Specific Integrated Circuit (ASIC), or the like. Theprocessor may be configured to run and/or execute application processesand/or other modules, processes and/or functions associated with thesystem and/or a network associated therewith. The underlying devicetechnologies may be provided in a variety of component types, e.g.,metal-oxide semiconductor field-effect transistor (MOSFET) technologieslike complementary metal-oxide semiconductor (CMOS), bipolartechnologies like emitter-coupled logic (ECL), polymer technologies(e.g., silicon-conjugated polymer and metal-conjugated polymer-metalstructures), mixed analog and digital, or the like.

In some embodiments, memory may include a database and may be, forexample, a random access memory (RAM), a memory buffer, a hard drive, anerasable programmable read-only memory (EPROM), an electrically erasableread-only memory (EEPROM), a read-only memory (ROM), Flash memory, etc.The memory may store instructions to cause the processor to executemodules, processes and/or functions associated with the system, such asone or more treatment plans, imaging data acquired during a previoustreatment session and/or current treatment session (e.g., real-timeimaging data), biological activity, physiological and/or anatomical dataextracted from imaging data, updated or adapted treatment plans, updatedor adapted dose delivery instructions, radiation therapy systeminstructions (e.g., that may direct the operation of the gantry,therapeutic radiation source, multi-leaf collimator, PET detectors,and/or any other components of a radiation therapy system), and imageand/or data processing associated with treatment delivery.

Some embodiments described herein relate to a computer storage productwith a non-transitory computer-readable medium (also may be referred toas a non-transitory processor-readable medium) having instructions orcomputer code thereon for performing various computer-implementedoperations. The computer-readable medium (or processor-readable medium)is non-transitory in the sense that it does not include transitorypropagating signals per se (e.g., a propagating electromagnetic wavecarrying information on a transmission medium such as space or a cable).The media and computer code (also may be referred to as code oralgorithm) may be those designed and constructed for the specificpurpose or purposes. Examples of non-transitory computer-readable mediainclude, but are not limited to, magnetic storage media such as harddisks, floppy disks, and magnetic tape; optical storage media such asCompact Disc/Digital Video Discs (CD/DVDs); Compact Disc-Read OnlyMemories (CD-ROMs), and holographic devices; magneto-optical storagemedia such as optical disks; solid state storage devices such as a solidstate drive (SSD) and a solid state hybrid drive (SSHD); carrier wavesignal processing modules; and hardware devices that are speciallyconfigured to store and execute program code, such asApplication-Specific Integrated Circuits (ASICs), Programmable LogicDevices (PLDs), Read-Only Memory (ROM), and Random-Access Memory (RAM)devices. Other embodiments described herein relate to a computer programproduct, which may include, for example, the instructions and/orcomputer code disclosed herein.

A user interface may serve as a communication interface between anoperator or clinician and the treatment planning system. The userinterface may comprise an input device and output device (e.g., touchscreen and display) and be configured to receive input data and outputdata from one or more of the support arm, external magnet, sensor,delivery device, input device, output device, network, database, andserver. Sensor data from one or more sensors may be received by userinterface and output visually, audibly, and/or through haptic feedbackby one or more output devices. As another example, operator control ofan input device (e.g., joystick, keyboard, touch screen) may be receivedby user and then processed by processor and memory for user interface tooutput a control signal to one or more support arms, external magnets,intracavity devices, and delivery devices.

In some variations, an output device may comprise a display deviceincluding at least one of a light emitting diode (LED), liquid crystaldisplay (LCD), electroluminescent display (ELD), plasma display panel(PDP), thin film transistor (TFT), organic light emitting diodes (OLED),electronic paper/e-ink display, laser display, and/or holographicdisplay.

In some variations, a radiation therapy system may be in communicationwith other computing devices via, for example, one or more networks,each of which may be any type of network (e.g., wired network, wirelessnetwork). A wireless network may refer to any type of digital networkthat is not connected by cables of any kind. Examples of wirelesscommunication in a wireless network include, but are not limited tocellular, radio, satellite, and microwave communication. However, awireless network may connect to a wired network in order to interfacewith the Internet, other carrier voice and data networks, businessnetworks, and personal networks. A wired network is typically carriedover copper twisted pair, coaxial cable and/or fiber optic cables. Thereare many different types of wired networks including wide area networks(WAN), metropolitan area networks (MAN), local area networks (LAN),Internet area networks (IAN), campus area networks (CAN), global areanetworks (GAN), like the Internet, and virtual private networks (VPN).Hereinafter, network refers to any combination of wireless, wired,public and private data networks that are typically interconnectedthrough the Internet, to provide a unified networking and informationaccess system.

Cellular communication may encompass technologies such as GSM, PCS, CDMAor GPRS, W-CDMA, EDGE or CDMA2000, LTE, WiMAX, and 5G networkingstandards. Some wireless network deployments combine networks frommultiple cellular networks or use a mix of cellular, Wi-Fi, andsatellite communication. In some embodiments, the systems, apparatuses,and methods described herein may include a radiofrequency receiver,transmitter, and/or optical (e.g., infrared) receiver and transmitter tocommunicate with one or more devices and/or networks.

1. A radiation therapy system comprising: a radiation source configured to direct one or more radiation pulses toward a PET-avid region of interest; a plurality of PET detectors, wherein the plurality of PET detectors are configured to detect positron annihilation photons; a current detector configured to measure a bias current of the plurality of PET detectors; and a controller configured to receive photon data output from the plurality of PET detectors, wherein the controller is configured to detect a pair of coincident positron annihilation photons by adjusting the photon data output using a gain factor calculated based on the measured bias current during a therapy session.
 2. The system of claim 1, wherein the controller is configured to adjust the gain factor when the bias current exceeds a threshold bias current value.
 3. The system of claim 1, wherein the gain factor is a ratio between the measured bias current and a magnitude of a photopeak shift of the detection of the positron annihilation photons in photon data output.
 4. The system of claim 1, wherein adjusting the photon data output comprises multiplying the photon data output by the gain factor or linearly shifting the photon data output by the gain factor.
 5. The system of claim 1, wherein the controller is configured to adjust the gain factor after a threshold number of radiation pulses have been directed toward the region of interest.
 6. The system of claim 5, wherein the threshold number of radiation pulses is approximately 1,000 radiation pulses.
 7. The system of claim 5, wherein the gain factor is a first gain factor and the threshold number of radiation pulses is a first threshold number of radiation pulses, and wherein the controller is configured to adjust the first gain factor to a second gain factor after a second threshold number of radiation pulses have been directed toward the region of interest.
 8. The system of claim 7, wherein the second gain factor is greater than the first gain factor and the second threshold number of radiation pulses is greater than the first threshold number of radiation pulses.
 9. The system of claim 1, wherein the controller is configured to calculate a photopeak location of annihilation photons based on the photon data output from the plurality of PET detectors and to adjust the gain factor based on shifts of the photopeak location from a baseline level.
 10. The system of claim 1, wherein the controller is configured to adjust the gain factor when a dark count rate of one or more of the plurality of PET detectors exceeds a threshold dark count rate.
 11. The system of claim 10, wherein the threshold dark count rate is from about 3 Mcps to about 10 Mcps.
 12. The system of claim 2, wherein the threshold bias current value is from about 0.1 mA to about 1 mA.
 13. The system of claim 1, wherein the controller is configured to adjust the gain factor when the amount of radiation emitted from the radiation source exceeds a threshold radiation level.
 14. The system of claim 13, wherein the threshold radiation level is from about 0.1 cGy/min to about 1 cGy/min.
 15. The system of claim 1, wherein the controller further comprises a signal processor and a switch configured to selectively communicate a PET detector output signal to the signal processor, wherein the switch is configured to suspend communication of the PET detector output signal to the signal processor for a predetermined period of time following each radiation pulse, wherein a ratio of the predetermined period of time to the duration of each radiation pulse is between about 25:1 to about 100:1.
 16. The system of claim 15, wherein the controller is configured to suspend communication of the PET detector output signal to the signal processor for the duration of each radiation pulse and the predetermined period of time following each radiation pulse.
 17. The system of claim 16, wherein the controller is configured to suspend communication of the PET detector output signal to the signal processor based on a gate signal.
 18. The system of claim 17, wherein the gate signal causes the controller to suspend communication of the PET detector output signal to the signal processor for 100 μs or more following each radiation pulse.
 19. The system of claim 18, wherein the gate signal causes the controller to suspend communication of the PET detector output signal to the signal processor for 200 μs or more following each radiation pulse.
 20. The system of claim 5, wherein the controller is configured to adjust the gain factor at least partially based on a timing schedule of the radiation pulses. 21-68. (canceled)
 69. A method for automatically adjusting a PET detector gain factor for detecting coincident positron annihilation photons, the method comprising: measuring a bias current of two or more PET detectors during a therapy session while a radiation source is activated; calculating a gain factor based on the measured bias current; and detecting a pair of coincident positron annihilation photons by adjusting photon data output from the two or more PET detectors by the calculated gain factor.
 70. The method of claim 69, further comprising determining whether the measured bias current meets or exceeds a threshold bias current value, and calculating the gain factor if the measured bias current meets or exceeds the threshold bias current value.
 71. The method of claim 69, wherein calculating the gain factor comprises calculating a ratio between the measured bias current and a magnitude of a photopeak shift of the detection of the coincident positron annihilation photons in photon data output.
 72. The method of claim 69, wherein adjusting the photon data output comprises multiplying the photon data output by the gain factor or linearly shifting the photon data output by the gain factor.
 73. The method of claim 69, further comprising determining whether the activated radiation source has applied a threshold number of radiation pulses toward a region of interest, and calculating the gain factor if the number of radiation pulses meets or exceeds the threshold number.
 74. The method of claim 73, wherein the threshold number of radiation pulses is approximately 1,000 radiation pulses.
 75. The method of claim 70, wherein the gain factor is a first gain factor, the measured bias current is a first bias current value, and the threshold bias current value is a first threshold bias current value, and wherein the method further comprises measuring a second bias current value, determining whether the second bias current value meets or exceeds a second threshold bias current value, and calculating a second gain factor based on the second bias current value if the second bias current value meets or exceeds the second threshold bias current value.
 76. The method of claim 75, wherein the second gain factor is greater than the first gain factor, and the second threshold bias current value is greater than the first threshold bias current value.
 77. The method of claim 69, further comprising calculating a photopeak location of positron annihilation photons based on the photon data output from the two or more PET detectors, and wherein the gain factor is calculated based on shifts of the photopeak location from a baseline level.
 78. The method of claim 69, further comprising determining whether a dark count rate of the two or more PET detectors meets or exceeds a threshold dark count rate, and calculating the gain factor if the dark count rate meets or exceeds the threshold dark count rate.
 79. The method of claim 78, wherein the threshold dark count rate is from about 3 Mcps to about 10 Mcps.
 80. The method of claim 70, wherein the threshold bias current value is from about 0.1 mA to about 1 mA.
 81. The method of claim 69, further comprising determining whether an amount of radiation emitted from the activated radiation source meets or exceeds a threshold radiation level, and calculating the gain factor if the amount of emitted radiation meets or exceeds the threshold radiation level.
 82. The method of claim 81, wherein the threshold radiation level is from about 0.1 cGy/min to about 1 cGy/min.
 83. The method of claim 73, wherein calculating the gain factor is based at least partially on a timing schedule of the radiation pulses emitted by the activated radiation source.
 84. The method of claim 69, wherein detecting the pair of coincident photon annihilation pair occurs 100 μs or more after the radiation source emits a radiation pulse. 